The treatment of bodily tissue with thermal energy to destroy it is useful for various therapeutic procedures. Thermal energy can be imparted to tissue using radio frequency electrical energy, microwave or lightwave electromagnetic energy, or ultrasonic vibrational energy. Each of these is absorbed by tissue and is converted to thermal energy or heat. Tissue can also be heated through conduction, in which a heated object is placed against tissue and from which the thermal energy is conducted into the tissue. One disadvantage associated with the application of these forms of energy is that the maximum heating often occurs at the interface between the therapeutic tool and the tissue. Radio frequency (RF) heating, for example, is effected by placing one or more electrodes against tissue to be treated and passing high frequency electrical current into the tissue. The current can flow between closely spaced emitting electrodes or between an emitting electrode and a larger, common electrode located remotely from the tissue to be heated. The maximum heating often occurs in the local tissue, immediately adjacent the emitting electrodes. This often causes the local tissue to desiccate, thereby reducing its electrical conductivity. As the tissue conductivity decreases, the impedance to current passing from the electrode to the tissue increases so that more voltage must be supplied to the electrodes to affect the surrounding, more distant tissue. The tissue temperature proximate to the electrode may approach 100.degree. C., so that water within the tissue boils to become water vapor. As this desiccation and/or vaporization process continues, the impedance of the local tissue rises so that a therapeutic level of current is not allowed to pass into the surrounding tissue.
Thus, conventional RF instruments are limited in the volume of tissue that can be treated, i.e., in the size of the thermal lesions they can create. Specifically, it is known that the level of tissue heating is proportional to the square of the electrical current density, and that the electrical current density in tissue generally falls as the square of the distance from the electrode. Therefore, the heating in tissue generally falls as the fourth power of distance from the electrode and the resulting tissue temperature therefore decreases rapidly as the distance from the electrode increases. This rapid fall-off in tissue temperature limits the volume of tissue that can be therapeutically treated using RF energy. While radio frequency (RF) electrodes can be repeatedly repositioned to treat additional tissue, the precise movement required for this task is time consuming.
The excessive heating that occurs around an RF electrode can be exacerbated by the use of metal as the emitting electrode. It is well-known that a voltage differential occurs when a metal contacts the saline solution in tissue and that current passing through a voltage differential causes energy to be dissipated at that interface. The voltage differential is a function of the specific metal and of the magnitude and direction of current flowing between the electrode and the tissue. Therefore, when the RF current passes through this voltage differential, an additional amount of energy is dissipated to produce a further increase in the temperature of the local tissue.
Attempts have been made to provide an RF device that increases lesion size. For example, existing RF instruments have been modified by increasing the electrode size or using multiple emitting electrodes. While such devices may provide some incremental increase in lesion size, larger tissue treatment volumes are desired. Other methods attempt to cool the tissue in the vicinity of the electrodes by circulating cool, i.e. body temperature (37.degree. C.) or lower than body temperature, water against or through the electrode and on into or against the nearby tissue being heated. Imran et al. describe one such device in U.S. Pat. No. 5,348,554. The device is a catheter with a metal RF electrode disposed at the distal end of the catheter. The catheter has a channel disposed within it, extending from the proximal end of the catheter to the electrode at the distal end. When the electrode is deployed against a body tissue, an RF electrical current is passed between the distal emitting electrode and a larger, counter electrode placed at some other location on the patient's body. Sterile saline solution is simultaneously passed through the catheter channel, which passes out of the catheter and against the tissue near the electrode. The flowing saline solution cools the tissue immediately adjacent to the electrode and moderates the excessive temperature this tissue would otherwise achieve. However, it does not significantly alter the temperature of the tissue even a short distance from the electrode, so that tissue can still overheat and limit the total size of the thermal lesion that can be created.
Mulier et al. describe a similar catheter in U.S. Pat. No. 5,431,649. The catheter includes a metal RF electrode at the distal end and an internal channel extending from the proximal end to the electrode. The catheter is deployed through the vasculature and into the heart. It is then inserted directly into the tissue to be treated and an RF current is passed between the catheter electrode and a larger, counter electrode remotely located on the patient's body. Cool sterile saline solution or body-temperature sterile saline solution is passed through the channel and into the tissue through the electrode. The saline solution cools the tissue most proximate to the electrode and convects the thermal energy it absorbs from this tissue into tissue located remotely from the electrode. This can create larger thermal lesions because of the cooling combined with the improved thermal transport in the tissue. However, this method can cause excessive cooling of the tissue and can actually decrease the size of the thermal lesion created or eliminate it entirely.
It is known that there is a time-temperature relationship in the thermal destruction of tissue. A threshold temperature for causing irreversible thermal damage to tissue is generally accepted to be about 41.degree. C. It is also known that the time required to achieve a particular level of cell necrosis decreases as the treatment temperature increases further above 41.degree. C. It is understood that the exact time/temperature relationship varies by cell type, but that there is a general relationship across many cell types that can be used to determine a desired thermal dose level. This relationship is commonly referred to as an equivalent time at 43.degree. C. expressed as EQU t.sub.eq,43.degree. C. =.intg.R.sup.(T(t)-43.degree.) dt (1)
where T is the tissue temperature, and R is a unitless indicator of therapeutic efficiency in a range between 0 and 5 (typically 2). Note that R is zero for temperatures below 41.degree. C. Thermal doses in the range of t.sub.eq,43.degree. C. =20 minutes to 1 hour are generally accepted as therapeutic. Thus, therapeutic temperature may refer to any temperature in excess of 41.degree. C., but the exact threshold is determined by the desired therapy duration and equation (1). For example, Nath, S. and Haines, D. E., Prog Card Dis 37(4):185-205 (1995) (Nath et al.) suggest a temperature of 50.degree. C. for one minute as therapeutic, which is an equivalent time at 43.degree. C. of 128 minutes with R=2. For maximum efficiency, the therapeutic temperature should be uniform throughout the tissue being treated so that the thermal dose is uniformly delivered.
Therefore, it is desirable to provide a device for the therapeutic treatment of tissue that creates a relatively large lesion size, eliminates overheating at the interface between the therapeutic tool and the tissue, and produces an approximately equal therapeutic temperature throughout the volume of tissue being treated.